The present invention relates to detection of radiation. More specifically, the present invention relates to detection of radiation, particularly X-ray radiation, using scanning.
Film has been used in medical X-ray imaging for more than 100 years, and is yet the dominating technique for X-ray detection at hospitals all over the world. The largest improvement came in the 1960s when embedding the film in a fluorescenting screen to increase the sensitivity drastically reduced the dose, but at the expense of a reduced position resolution.
In the middle of the 1990s, the first digital technique appeared on the market, where the X-ray illumination gives digital signals in the detector, for creating digital images. These detectors either convert the X-ray flux to visible photons in a scintillator or to charge in a semiconductor. The light from the scintillator is detected with e.g. a TFT or a CCD (Charged Coupled Device), and the charge from the semiconductor is detected using e.g. a TFT (Thin Film Transistor). These techniques solves some of the problems associated with using film, but still have some drawbacks.
The basic CCD consists of a series of metal oxide-semiconductor capacitors that are fabricated close together on a semiconductor surface. Today, CCDs are used in a wide variety of indirect-conversion X-ray imaging devices, including large-area radiographic imaging systems. The single most salient characteristic of CCDs with regard to radiography is that they are physically small in size, typically 2-4 cm2, which is much smaller than typically projected X-ray areas. Because of this, cost-effective CCD-based radiographic systems must include some means of optical coupling to reduce the size of the projected visible light image and to transfer the image to the face of one or more CCDs. Some systems are based on an array of CCD cameras, each of which is coupled to a scintillator by a lens or fibre optic taper.
Recent advantages in photolithography and electronic micro fabrication techniques have enabled the development of large area x-ray detectors with integrated readout mechanisms based on arrays of TFT. Unlike CCD-based detectors that require optical coupling and image demagnification, TFT-based, flat panel systems are constructed such that the pixel charge collection and readout electronics for each pixel are immediately adjacent to the position of the X-ray interactions. TFT-arrays are used as active electronic elements in both indirect and direct conversion flat panel detectors. In indirect systems X-rays produces light in scintillators, which is optically coupled to a light sensitive device where the produced light is converted into charges, which are detected. In direct systems X-rays directly produces charges, which are detected.
TFT-arrays are typically deposited onto a glass substrate in multiple layers, beginning with readout electronics at the lowest level and followed by charged collector arrays at higher levels. Depending on the type of detector, X-ray-elements (direct conversion), light-sensitive elements (indirect conversion) or both, are deposited to form the top layer of the electronic sandwich structure.
Typically, amorphous selenium is used as conversion material, due to its good X-ray detection possibilities and high intrinsic spatial resolution, in direct conversion TFT systems. Before exposure, an electric field is applied across the amorphous selenium layer through a bias electrode on the top surface of the selenium. As X-rays are absorbed in the detector, electrons and holes are released within the selenium, and due to the electric field within the selenium, electric charges are drawn directly to the charge collecting electrodes. Pixels are separated by means of field shaping within the selenium layer.
Indirect conversion systems based on TFT-arrays are constructed by adding an amorphous silicon photodiode circuitry and a scintillator optically coupled to the top layers of the TFT sandwich. When X-rays strike the scintillator, visible light is emitted proportional to the incident X-ray energy. Visible light photons are then converted into an electric charge by the photodiode array.
TFT-arrays, however, have a limited position resolution set by the minimum pixel size, which is currently limited to approximately 100 xcexcmxc3x97100 xcexcm. Researchers are currently reaching minimum pixel sizes as small as 70 xcexcmxc3x9770 xcexcm, whereas for affordable consumer products pixel sizes are still reasonably large, for instance approximately 450 xcexcmxc3x97450 xcexcm in computer screens. It would be advantageous if the resolution achieved in X-ray imaging could be made even greater.
Large TFT-array systems are also costly. An array, with maximum resolution, i.e. 100 xcexcmxc3x97100 xcexcm, and a size of 25 cmxc3x9725 cm will have 6.25 million individual pixels and would cost more than approximately $100000. The high cost is mainly due to a low yield when trying to minimize the area of the thin-film transistor in amorphous silicon under the requirement that very few dead pixels are allowed.
One individual pixel comprises for indirect systems, apart from the X-ray detecting means, a photo diode for detecting the light, a capacitor for storing the charge, a switch transistor for reading out the value of the pixel, data and address lines and an isolation distance. This means that a relative small part is available for the X-ray detecting means, for a 100 xcexcmxc3x97100 xcexcm pixel only approximately 50%. This is called the fill factor.
It would be advantageous if the cost could be reduced, while keeping the position resolution high and the size large.
Another problem with conventional X-ray detecting apparatuses is that the X-rays is scattered in the object to be imaged. Approximately as much as 50% of the incident X-ray radiation is randomly scattered, producing a fog or haze in the X-ray image.
It would be advantageous if the amount of scattered X-ray radiation could be reduced.
It is a main object of the present invention to provide such apparatus and method, wherein the obtained resolution in X-ray imaging is increased.
It is in this respect a particular object of the invention to provide such apparatus and method that allow for X-ray imaging of large objects with improved resolution.
It is still a further object of the invention to provide such apparatus and method that reduce the administrated radiation dose to an object, such as an examined patient.
It is still a further object of the invention to provide such apparatus and method that reduces the noise in the image caused by scattered X-ray radiation.
It is still a further object of the invention to provide such apparatus and method that reduces the cost of X-ray detectors.
It is still a further object of the invention to provide such apparatus and method that reduces the number of electronic channels in large area X-ray detectors.
These objects among others are attained, according to the present invention, by apparatus and methods as claimed in the appended patent claims.
An advantage of the present invention is that the resolution of the X-ray image is increased.
A further advantage is that the administrated radiation dose to an object, to be imaged, is reduced.
A further advantage is that the amount of scattered X-ray radiation in the object to be imaged is reduced.
Further characteristics of the invention and advantages thereof will be evident from the following detailed description of embodiments of the invention.